Spinal Cord Stimulator System

ABSTRACT

Spinal cord stimulation (SCS) system having a recharging system with self alignment, a system for mapping current fields using a completely wireless system, multiple independent electrode stimulation outsource, and IPG control through software on Smartphone/mobile device and tablet hardware during trial and permanent implants. SCS system can include multiple electrodes, multiple, independently programmable, stimulation channels within an implantable pulse generator (IPG) providing concurrent, but unique stimulation fields. SCS system can include a replenishable power source, rechargeable using transcutaneous power transmissions between antenna coil pairs. An external charger unit, having its own rechargeable battery, can charge the IPG replenishable power source. A real-time clock can provide an auto-run schedule for daily stimulation. A bi-directional telemetry link informs the patient or clinician the status of the system, including the state of charge of the IPG battery. Other processing circuitry in current IPG allows electrode impedance measurements to be made.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a non-provisional application that claims priorityto provisional application No. 61/792,654 filed on Mar. 15, 2013, whichis incorporated in its entirety herein.

TECHNICAL FIELD

This disclosure relates to stimulators using electrical pulses in amedical context, and more particularly, applying electrical pulsestimulators to the spinal cord to control pain.

BACKGROUND

A Spinal Cord Stimulator (SCS) is used to exert pulsed electricalsignals to the spinal cord to control chronic pain. Spinal cordstimulation, in its simplest form, comprises stimulating electrodesimplanted in the epidural space, an electrical pulse generator implantedin the lower abdominal area or gluteal region, conducting wiresconnecting the electrodes to the electrical pulse generator, anelectrical pulse generator remote control, and an electrical pulsegenerator charger. Spinal cord stimulation has notable analgesicproperties and, at the present, is used mostly in the treatment offailed back surgery syndrome, complex regional pain syndrome andrefractory pain due to ischemia.

Electrotherapy of pain by neurostimulation began shortly after Melzackand Wall proposed the gate control theory in 1965. This theory proposedthat nerves carrying painful peripheral stimuli and nerves carryingtouch and vibratory sensation both terminate in the dorsal horn (thegate) of the spinal cord. It was hypothesized that input to the dorsalhorn of the spinal cord could be manipulated to “close the gate” to thenerves. As an application of the gate control theory, Shealy et al.implanted the first spinal cord stimulator device directly on the dorsalcolumn for the treatment of chronic pain in 1971.

Spinal cord stimulation does not eliminate pain. The electrical impulsesfrom the stimulator override the pain messages so that the patient doesnot feel the pain intensely. In essence, the stimulator masks the pain.A trial implantation is performed before implanting the permanentstimulator. The physician first implants a trial stimulator through theskin (percutaneously) perform stimulations as a trial run. Because apercutaneous trial stimulator tends to move from its original location,it is considered temporary. If the trial is successful, the physiciancan then implant a permanent stimulator. The permanent stimulator isimplanted under the skin of the abdomen, and with the leads insertedunder the skin and subcutaneously fed to and inserted into the spinalcanal. This placement of the stimulator in the abdomen is a more stable,effective location. The leads, which consist of an array of electrodes,can be percutaneous type or paddle type. Percutaneous electrodes areeasier to insert in comparison with paddle type, which are inserted viaincision over spinal cord and laminectomy.

There are a number of the problems that exist in currently available SCSsystems that limit the full benefits of dorsal column stimulation froman effectiveness and patient user friendly perspective. One problem isthat current SCS systems are limited to only 16 electrodes with amaximum of 16 independent current sources. Another problem is thatcurrent SCS systems have complicated trialing methods that involvemultiple gadgets and hardware. Another problem is that patients mustcarry an independent remote control in order to control the IPG in theirdaily lives.

SUMMARY

Disclosed are the following features included within a spinal cordstimulation system: (1) a recharging system with self alignment, (2) asystem for mapping current fields using a completely wireless system,(3) multiple independent electrode stimulation outsource, and (4) IPGcontrol, during trial and permanent implants, through software ongeneric Smartphone/mobile device and tablet hardware. The SCS system caninclude multiple electrodes, and multiple, independently programmable,stimulation channels within an implantable pulse generator (IPG) wherethe channels can provide concurrent, but unique stimulation fields,permitting virtual electrodes to be realized. The SCS system can includea replenishable power source (e.g., rechargeable battery) that may berecharged using transcutaneous power transmissions between antenna coilpairs. An external charger unit, having a rechargeable battery can beused to charge the IPG replenishable power source. A real-time clock canprovide an auto-run schedule for daily stimulation. An includedbi-directional telemetry link in the system can inform the patient orclinician of the status of the system, including the state of charge ofthe IPG battery. Other processing circuitry in the IPG allows electrodeimpedance measurements to be made. Circuitry provided in the externalbattery charger can provide alignment detection for the coil pairs. FIG.1 depicts a SCS system, as described herein, for use during the trialperiod and the permanent implantation.

The SCS system, as disclosed herein, is superior to existing SCS systemsin that the SCS system, as disclosed herein, can provide a stimulus to aselected pair or group of a multiplicity of electrodes, e.g., 32electrodes, grouped into multiple channels, e.g., 6 channels. In anembodiment, each electrode is able to produce a programmable constantoutput current of at least 12 mA over a range of output voltages thatmay go as high as 16 volts. In another embodiment, the implant portionof the SCS system includes a rechargeable power source, e.g., one ormore rechargeable batteries. The SCS system described herein requiresonly an occasional recharge, has an implanted portion smaller thanexisting implant systems, has a self aligning feature to guide thepatient in aligning the charger over the implanted IPG for the mostefficient power recharge, has a life of at least 10 years at typicalsettings, offers a simple connection scheme for detachably connecting alead system thereto, and is extremely reliable.

In an embodiment, each of the electrodes included within the stimuluschannels can deliver up to 12.7 mA of current over the entire range ofoutput voltages, and can be combined with other electrodes to delivercurrent up to a maximum of 20 mA. Additionally, the SCS system providesthe ability to stimulate simultaneously on all available electrodes inthe SCS system. That is, in operation, each electrode can be groupedwith at least one additional electrode to form one channel. The SCSsystem allows the activation of electrodes to at least 10 channels. Inone embodiment, such grouping is achieved by a low impedance switchingmatrix that allows any electrode contact or the system case (which maybe used as a common, or indifferent, electrode) to be connected to anyother electrode. In another embodiment, programmable output currentDAC's (digital-to-analog converters) are connected to each electrodenode, so that, when enabled, any electrode node can be grouped with anyother electrode node that is enabled at the same time, therebyeliminating the need for a low impedance switching matrix. Thisadvantageous feature allows the clinician to provide unique electricalstimulation fields for each current channel, heretofore unavailable withother “multichannel” stimulation systems (which “multi-channel”stimulation systems are really multiplexed single channel stimulationsystems). Moreover, this feature, combined with multicontact electrodesarranged in two or three dimensional arrays, allows “virtual electrodes”to be realized, where a “virtual electrode” comprises an electrode thatappears to be at a certain physical location, but in actuality is notphysically located at the certain physical location. Rather, the“virtual electrode” results from the vector combination of electricalfields from two or more electrodes that are activated simultaneously.

In embodiments, the SCS system includes an implantable pulse generator(IPG) powered by a rechargeable internal battery, e.g., a rechargeablelithium ion battery providing an output voltage that varies from about4.1 volts, when fully charged, to about 3.5 volts.

Embodiments are comprised of components previously not provided on SCSsystems. The components are comprised of a number of differentsub-components, as described herein. The SCS system can be comprised ofan permanent implantable IPG, an implantable trial IPG, a wirelessdongle, an IPG charger, clinical programmer software, patient programmersoftware, leads (percutaneous and paddle), lead anchors, lead splitters,lead extensions, and accessories. FIG. 1 depicts the components duringtrial and permanent implantation.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts various components that can be included in a spinal cordstimulation system, according to an embodiment.

FIG. 2 depicts an exploded view of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 3 depicts a feedthrough assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 4 depicts a lead contact system of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 5 depicts a lead contact assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 6 depicts a head unit assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 7 depicts an RF antenna of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 8 depicts a percutaneous lead, according to an embodiment.

FIG. 9 depicts a paddle lead, according to an embodiment.

FIG. 10 depicts a lead extension, according to an embodiment.

FIG. 11 depicts a lead splitter, according to an embodiment.

FIG. 12 depicts a sleeve anchor, according to an embodiment.

FIG. 13 depicts a mechanical locking anchor, according to an embodiment.

FIG. 14 illustrates communication via a wireless dongle with atablet/clinician programmer and smartphone/mobile/patient programmerduring trial and/or permanent implantation, according to an embodiment.

FIG. 15 depicts a Tuohy needle, according to an embodiment.

FIG. 16 depicts a stylet, according to an embodiment.

FIG. 17 depicts a passing elevator, according to an embodiment.

FIG. 18 depicts a tunneling tool, according to an embodiment.

FIG. 19 depicts a torque wrench, according to an embodiment.

DETAILED DESCRIPTION Implantable Pulse Generator (IPG)

FIG. 1 illustrates various components that can be included in a SCSsystem for the trial and the permanent installation periods. The spinalcord stimulator (SCS) 100 is an implantable device used to deliverelectrical pulse therapy to the spinal cord in order to treat chronicpain. The implantable components of the system consist of an ImplantablePulse Generator (IPG) 102 and a multitude of stimulation electrodes 130.The IPG 102 is implanted subcutaneously, no more than 30 mm deep in anarea that is comfortable for the patient while the stimulationelectrodes 130 are implanted directly in the epidural space. Theelectrodes 130 are wired to the IPG 102 via leads 140, 141 which keepthe stimulation pulses isolated from each other in order to deliver thecorrect therapy to each individual electrode 130.

The therapy delivered consists of electrical pulses with controlledcurrent amplitude ranging from +12.7 to −12.7 mA (current range 0-25.4mA). These pulses can be programmed in both length and frequency from 10μS to 20000 μS and 0.5 Hz to 1200 Hz. At any given moment, the sum ofthe currents sourced from the anodic electrodes 130 must equal the sumof the currents sunk by the cathodic electrodes 130. In addition, eachindividual pulse is bi-phasic, meaning that once the initial pulsefinishes another pulse of opposite amplitude is generated after a setholdoff period. The electrodes 130 may be grouped into stimulation setsin order to deliver the pulses over a wider area or to target specificareas, but the sum of the currents being sourced at any one given timemay not exceed 20 mA. A user can also program different stimulation sets(up to eight) with different parameters in order to target differentareas with different therapies.

FIG. 2 depicts an exploded view of an IPG 102. The IPG 102 consists oftwo major active components 104, 106, a battery 108, antenna 110, somesupport circuitry, and a multitude of output capacitors 112. The firstof the major active components is the microcontroller 104 transceiver104. It is responsible for receiving, decoding, and execution bothcommands and requests from the external remote. If necessary it passesthese commands or requests onto the second major component, the ASIC106. The ASIC 106 receives the digital data from the microcontroller 104and performs the entire signal processing to generate the signalsnecessary for stimulation. These signals are then passed onto thestimulation electrodes 130 in the epidural space.

The ASIC 106 is comprised of a digital section and an analog section.The digital section is divided into multiple sections including; TimingGenerators, Arbitration Control, Pulse Burst Conditioner, and ElectrodeLogic. The analog section receives the incoming pulses from the digitalsection and amplifies them in order to deliver the correct therapy.There are also a multitude of digital register memory elements that eachsection utilizes, both digital and analog.

The digital elements in the ASIC 106 are all made up of standard subsetsof digital logic including logic gates, timers, counters, registers,comparators, flip-flips, and decoders. These elements are ideal forprocessing the stimulation pulses as all of them can function extremelyfast—orders of magnitudes faster than the required pulse width. The onedrawback is that they must all function at one single voltage, usually5.0, 3.3, 2.5, or 1.8 volts. Therefore they are not suitable for thefinal stage in which the pulses are amplified in order to deliver theconstant current pulses.

The timing generators are the base of each of the stimulation sets. Itgenerates the actual rising and falling edge triggers for each phase ofthe bi-phasic pulse. It accomplishes this by taking the incoming clockthat is fed from the microcontroller 104 and feeding it into a counter.For the purpose of this discussion, assume the counter simply countsthese rising clock edges infinitely. The output of the counter is fedinto six different comparators. The comparators other input is connectedto specific registers that are programmed by the microcontroller 104.When the count equals the value stored in the register, the comparatorasserts a positive signal.

The first comparator is connected to the SET signal of a SR flip flop.The SR flip flop stays positive until the RESET signal is asserted,which the second comparator is connected to. The output of the SR flipflop is the first phase of the bi-phasic pulse. Its rising & fallingedges are values stored in the registers and programmed by themicrocontroller 104. The third and fourth comparators & registers workin exactly the same way to produce the second phase of the bi-phasicpulse using the second SR flip flop.

The fifth comparator is connected the RESET of the final SR-Flip flop inthe timing generator. This flip flop is SET by the first comparator,which is the rising edge of the first pulse. The RESET is then triggeredby the value the microprocessor programmed into the register connectedto the comparator. This allows for a ‘holdoff’ period after the fallingedge of the second pulse. The output of this third SR flip flop can bethought of as an envelope of the biphasic pulses indicating when thisparticular timing generator is active.

The final comparator of the system is once again connected to a registerthat stores the frequency values from the microprocessor. Essentiallywhen the count reaches this value it triggers the comparator which isfed back to the counter to reset it to zero and beginning the entirepulse generation cycle again. The ASIC 106 may contain many of thesetiming generators as each can control anywhere from two to all of theelectrodes 130 connected to the IPG 102 at a time. However, when thereis more than one timing generator and multiple channels have beenactively programmed then there needs to be a mechanism for suppressing asecond channel from turning on when another is already active.

This brings us to the next circuit block contained in the IPG 102, thearbitrator. The arbitrator functions by looking at each of the timinggenerators' envelope signals and makes sure only one can be active at atime. If a second tries to activate then the arbitrator suppresses thatsignal.

It accomplishes this by bringing each of the channel envelope signalsinto a rising edge detection circuit. Once one is triggered it is fedinto the SET pin of an SR flip flop. The output of this SR-flip flop isfed into all of the other rising edge detectors in order to suppressthem from triggering. The channel envelope signal is also fed into afalling-edge detector which is then fed into the RESET of the same SRflip flop. The output of the SR flip flops are then connected toswitches whose outputs are all tied together that turn on/off thatchannels particular biphasic pulse train. Therefore the output of thiscircuit element is a single bi-phasic pulse train and a signaldesignating which timing generator that particular pulse train issourced from. Essentially, the circuit looks for a channel to go active.Once it finds one it suppresses all others until that channel becomesinactive.

The next section of the circuit works very similarly to the timinggenerators to create a high speed burst pulse train that is thencombined with the stimulation pulse train to create a bursted bi-phasicpulse train if desired.

It accomplishes this by taking the incoming clock that is fed from themicrocontroller 104 and feeding it into a counter. For the purpose ofthis discussion, assume the counter simply counts these rising clockedges infinitely. The counter is only active when during a single phaseof the bi-phasic signal and begins counting as soon as the rising edgeis detected. The output of the counter is fed into a comparator, alongwith a microcontroller-programmed register, whose output is connected tothe reset pin on the counter. Therefore this counter will simply countto a programmed value & reset. This programmed value is the burstfrequency.

The output of the comparator is then fed into an edge detection circuitand then a flip flop that combines it with the actual stimulation pulsetrain to create a single phase bursted stimulation pulse. The entirecircuit is duplicated for the second phase of the signal resulting inthe desired bursted bi-phasic pulse train. The stimulation signal is nowready to be handed over to the electrode logic stage.

The electrode logic conditions and directs the bi-phasic signals to theanalog section of the ASIC 106. At this point, the bi-phasic signalscontain all of the pertinent timing information, but none of therequired amplitude information. The incoming signals include thebi-phasic pulse train and another signal designating which timinggenerator the current active train came from. Each electrode logic cellhas a register for each timing generator that stores this particularelectrode's 130 amplitude values for that timing generator. Theelectrode logic cell uses the designation signal to determine whichregister to pull the amplitude values from, e.g. if the third timinggenerator is passed through the arbitration circuit then the electrodelogic would read the value from the third register.

Once the value is pulled from the register, it goes through a series oflogic gates. The gates first determine that the electrode 130 should beactive. If not, they proceed no further and do not activate the analogsection of the electrode output, thereby saving precious battery 108power. Next they determine if this particular electrode 130 is an anodeor cathode. If it is deemed to be an anode, the electrode logic passesthe amplitude information and the biphasic signal to the positivecurrent (digital to analog converter) DAC in the analog section of theASIC 106. If it is deemed to be a cathode, the electrode logic passesthe amplitude information and the biphasic signal to the negativecurrent DAC in the analog section of the ASIC 106. The electrode logiccircuit must make these decisions for each phase of the bi-phasic signalas every electrode 130 will switch between being an anode and a cathode.

The analog elements in the ASIC 106 are uniquely designed in order toproduce the desired signals. The basis of analog IC design is the fieldeffect transistor (FET) and the type of high current multiple outputdesign required in SCS 100 means that the bulk of the silicon in theASIC 106 will be dedicated to the analog section.

The signals from the electrode output are fed into each current DAC whenthat specific electrode 130 should be activated. Each electrode 130 hasa positive and a negative current DAC, triggered by the electrode logicand both are never active at the same time. The job of each current DACis, when activated, to take the digital value representing a stimulationcurrent amplitude and produce an analog representation of this value tobe fed into the output stage. This circuit forms half of the barrierbetween the digital and analog sections of the ASIC 106.

The digital section of the ASIC 106 is built upon a technology that onlyallows small voltages to exist. In moving to the analog section, theoutput of the current DAC (which is a low level analog signal) must beamplified to a higher voltage for use in the analog section. The circuitthat performs this task is called a power level shifter. Because thiscircuit is built upon two different manufacturing technologies andrequires high precision analog circuits built upon a digital base, it isextremely difficult to implement.

Once the voltages have been converted for usage in the analog portion ofthe ASIC 106 they are passed on to the output current stages. There aretwo current sources per electrode output. One will source a positivecurrent and one will sink a negative current, but they will never bothbe active simultaneously. The current sources themselves are made up ofanalog elements similar to a Howland current source. There is an inputstage, and amplification stage with feedback through a sensing componentto maintain the constant current. The input stage takes the analogvoltage values from the power level shifter and produces an output pulsedesignated for the amplifier. The amplifier then creates the pulses ofvarying voltages but constant current flow. The sources are capable ofsourcing or sinking up to 12.7 mA at 0.1 mA resolution into a load of upto 1.2 k Ohms. This translates into range of 15 volts, which will varydepending on the load in order to keep the current constant.

The microcontroller 104 to ASIC 106 interface is designed to be assimple as possible with minimal bus ‘chatter’ in order to save battery108 life. The ASIC 106 will essentially be a collection of registersprogrammed via a standard I²C or SPI bus. Since the ASIC 106 is handlingall the power management, there will also be a power good (PG) linebetween the two chips 104, 106 in order to let the microcontroller 104know when it is safe to power up. The ASIC 106 will also need to use apin on the microcontroller 104 in order to generate a hardware interruptin case anything goes awry in the ASIC 106. The final connection is thetime base for all of the stimulation circuitry. The ASIC 106 willrequire two clocks, one for its internal digital circuitry which will befed directly from the microcontroller 104 clock output, and one to baseall stimulation off of which will need to be synthesized by themicrocontroller 104 and fed to the ASIC 106. All commands and requeststo the ASIC 106 will be made over the I²C or SPI bus and will involvesimply reading a register address or writing to a register. Even whenthe ASIC 106 generates a hardware interrupt, it will be theresponsibility of the microcontroller 104 to poll the ASIC 106 anddetermine the cause of the interrupt.

The wireless interface is based upon the FCCs MedRadio standardoperating in the 402-405 Mhz range utilizing up to 10 channels fortelemetry. The protocol is envisioned to be very simple once again inorder to minimize transmission and maximize battery 108 life. Allprocessing will take place on the user remote/programmer and the onlydata transmitted is exactly what will be used in the microcontroller 104to ASIC 106 bus. That is, all of the wireless packets will containnecessary overhead information along with only a register address, datato store in the register, and a command byte instructing themicrocontroller 104 what to do with the data. The overhead section ofthe wireless protocol will contain synchronization bits, start bytes, anaddress which is synchronized with the IPG's 102 serial number, and aCRC byte to assure proper transmission. It is essential to keep thepacket length as small as possible in order to maintain battery 108life. Since the IPG 102 cannot listen for packets all the time due tobattery 108 life, it cycles on for a duty cycle of less than 0.05% ofthe time. This time value can be kept small as long as the data packetsare also small. The user commands needed to run the system are executedby the entire system using flows.

The IPG 102 uses an implantable grade Li ion battery 108 with 215 mAHrwith zero volt technology. The voltage of the battery 108 at fullcapacity is 4.1 V and it supplies current only until it is drained up to3.3 V which is considered as 100% discharged. The remaining capacity ofthe battery 108 can be estimated at any time by measuring the voltageacross the terminals. The maximum charge rate is 107.5 mA. A ConstantCurrent, Constant Voltage (CCCV) type of regulation can be applied forfaster charging of the battery 108.

The internal secondary coil 109 is made up of 30 turns of 30 AWG coppermagnet wires. The ID, OD, and the thickness of the coil are 30, 32, and2 mm, respectively. Inductance L2 is measured to be 58 uH, a 80 nFcapacitor is connected to it to make a series resonance tank at 74 kHzfrequency. In the art of induction charging, two types of rectifiers areconsidered to convert the induced AC into usable DC, either a bridgefull wave rectifier or a voltage doubler full wave rectifier. To obtaina higher voltage, the voltage double full wave rectifier is used in thisapplication. The rectifier is built with high speed Schottky diodes toimprove its function at high frequencies of the order 100 kH. A Zenerdiode and also a 5V voltage regulator are used for regulation. Thiscircuit will be able to induce AC voltage, rectify to DC, regulate to 5Vand supply 100 mA current to power management IC that charges theinternal battery 108 by CCCV regulation.

The regulated 5V 100 mA output from the resonance tank is fed to, forexample, a Power Management Integrated Circuit (PMIC) MCP73843. Thisparticular chip was specially designed by Microchip to charge a Li ionbattery 108 to 4.1 V by CCCV regulation. The fast charge current can beregulated by changing a resistor; it is set to threshold current of 96mA in this circuit. The chip charges the battery 108 to 4.1V as long asit receives current more than 96 mA. However, if the supply currentdrops below 96 mA, it stops to charge the battery 108 until the supplyis higher than 96 again. For various practical reasons, if the distancebetween the coils increases, the internal secondary coil 109 receiveslesser current than the regulated value, and instead of charging thebattery 108 slowly, it pauses the charging completely until it receivesmore than 96 mA. It is understood to those with skill in the art thatother power management chips can be used and the power management chipis not limited to the PMIC MCP738432 chip.

All the functions of the IPG 102 are controlled from outside using ahand held remote controller specially designed for this device. Alongwith the remote control, an additional control is desirable to operatethe IPG 102 if the remote control was lost or damaged. For this purposea Hall effect based magnet switch was incorporated to either turn ON orturn OFF the IPG 102 using an external piece of magnet. Magnet switchacts as a master control for the IPG 102 to turn on or off. A south poleof sufficient strength turns the output on and a north pole ofsufficient strength is necessary to turn the output off. The output islatched so that the switch continues to hold the state even after themagnet is removed from its vicinity.

The IPG 102 is an active medical implant that generates an electricalsignal that stimulates the spinal cord. The signal is carried through astimulation lead 140 that plugs directly into the IPG 102. The IPG 102recharges wirelessly through an induction coil 109, and communicates viaRF radio antenna 110 to change stimulation parameters. The IPG 102 isimplanted up to 3 cm below the surface of the skin and is fixed to thefascia by passing two sutures through holes in the epoxy header 114. Theleads 140 are electrically connected to the IPG 102 through a leadcontact system 116, a cylindrical spring-based contact system withinter-contact silicone seals. The leads 140 are secured to the IPG 102with a set screw 117 that actuates within locking housing 118. Set screwcompression on the lead's 140 fixation contact is governed by adisposable torque wrench. The wireless recharging is achieved byaligning the exterior induction coil on the charger with the internalinduction coil 109 within the IPG 102. The RF antenna within theremote's dongle 200 communicates with the RF antenna 110 in the IPG's102 epoxy header 114. FIG. 2 illustrates an exploded view of the IPG 102assembly.

The IPG 102 is an assembly of a hermetic titanium (6Al-4V) casing 120which houses the battery 108, circuitry 104, 106, and charging coil 109,with an epoxy header 114, which houses the lead contact assembly 116,locking housing 118, and RF antenna 110. The internal electronics areconnected to the components within the epoxy head through a hermeticfeedthrough 122, as shown in FIG. 3. The feedthrough 122 is a titanium(6Al-4V) flange with an alumina window and gold trimming. Within thealumina window are thirty-four platinum-iridium (90-10) pins thatinterface internally with a direct solder to the circuit board, andexternally with a series of platinum iridium wires laser-welded to theantenna 110 and lead contacts 126. The IPG 102 has the ability tointerface with 32 electrical contacts 126, which are arranged in fourrows of eight contacts 126. Thirty two of the feedthrough's 122 pins 124will interface with the contacts 126, while two will interface with theantenna 110, one to the ground plane and one to the antenna 110 feed.

FIGS. 4 and 5 depict a lead contact system 115 and assembly 116,respectively. The lead contacts 126 consist of an MP35N housing 128 witha platinum-iridium 90-10 spring 129. Each contact 126 is separated by asilicone seal 127. At the proximal end of each stack of 8 contacts 126is a titanium (6Al-4V) cap 125 which acts as a stop for the lead 140. Atthe distal end is a titanium (6Al-4V) set screw 119 and block 118 forlead fixation. At the lead entrance point there is a silicone tube 123which provides strain relief as the lead 140 exits the head unit 114,and above the set screw 119 is another silicone tube 131 with a smallinternal canal which allows the torque wrench to enter but does notallow the set screw 119 to back out. In addition to the contacts 126 andantenna 110, the header 114 also contains a radiopaque titanium (6Al-4V)tag 132 which allows for identification of the device under fluoroscopy.The overmold of the header 114 is Epotek 301, a two-part, biocompatibleepoxy. FIGS. 4, 5, 6, and 7 depict illustrations of lead contact system115, lead contact assembly 116, head unit assembly 114, and RF antenna110, respectively.

Internal to the titanium (6Al-4V) case 120 are the circuit board 105,battery 108, charging coil 109, and internal plastic support frame. Thecircuit board 105 will be a multi-layered FR-4 board with copper tracesand solder mask coating. Non-solder masked areas of the board will beelectroless nickel immersion gold. The implantable battery 108, allsurface mount components, ASIC 106, microcontroller 104, charging coil109, and feedthrough 122 will be soldered to the circuit board 105. Theplastic frame, made of either polycarbonate or ABS, will maintain thebattery's 108 position and provide a snug fit between the circuitry 105and case 120 to prevent movement. The charging coil 109 is a woundcoated copper.

Leads

The percutaneous stimulation leads 140, as depicted in FIG. 8, are afully implantable electrical medical accessory to be used in conjunctionwith the implantable SCS 100. The primary function of the lead is tocarry electrical signals from the IPG 102 to the target stimulation areaon the spinal cord. Percutaneous stimulation leads 140 providecircumferential stimulation. The percutaneous stimulation leads 140 mustprovide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and stimulation area. The leads 140 are surgicallyimplanted through a spinal needle, or epidural needle, and are driventhrough the spinal canal using a steering stylet that passes through thecenter of the lead 140. The leads 140 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 140. The leads 140 are secured at theproximal end with a set-screw 119 on the IPG 102 which applies radialpressure to a blank contact on the distal end of the proximal contacts.

The percutaneous stimulation leads 140 consist of a combination ofimplantable materials. Stimulation electrodes 130 at the distal end andelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy. This alloy is utilized for its bio-compatibilityand electrical conductivity. The electrodes 130 are geometricallycylindrical. The polymeric body of the lead 140 is polyurethane, whichis chosen for its bio-compatibility, flexibility, and high lubricity todecrease friction while being passed through tissue. The polyurethanetubing has a multi-lumen cross section, with one center lumen 142 andeight outer lumens 144. The center lumen 142 acts as a canal to containthe steering stylet during implantation, while the outer lumens 144provide electrical and mechanical separation between the wires 146 thatcarry stimulation from the proximal contacts to distal electrodes 130.These wires 146 are a bundle of MP35N strands with a 28% silver core.The wires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. The wires 146are laser welded to the contacts and electrodes 130, creating anelectrical connection between respective contacts on the proximal anddistal ends. The leads 140 employ a platinum-iridium plug 148, moldedinto the distal tip of the center lumen 142 to prevent the tip of thesteering stylet from puncturing the distal tip of the lead 140. Leads140 are available in a variety of 4 and 8 electrode 130 configurations.These leads 140 have 4 and 8 proximal contacts (+1 fixation contact),respectively. Configurations vary by electrode 130 number, electrode 130spacing, electrode 130 length, and overall lead 140 length.

The paddle stimulation leads 141, as depicted in FIG. 9, are a fullyimplantable electrical medical accessory to be used in conjunction withthe implantable SCS 100. The primary function of the paddle lead 141 isto carry electrical signals from the IPG 102 to the target stimulationarea on the spinal cord. The paddle leads 141 provide uni-directionstimulation across a 2-dimensional array of electrodes 130, allowing forgreater precision in targeting stimulation zones. The paddle stimulationleads 141 must provide a robust, flexible, and bio-compatible electricconnection between the IPG 102 and stimulation area. The leads 141 aresurgically implanted through a small incision, usually in conjunctionwith a laminotomy or laminectomy, and are positioned using forceps or asimilar surgical tool. The leads 141 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 141. The leads 141 are secured at theproximal end with a set-screw on the IPG 102 which applies radialpressure to a fixation contact on the distal end of the proximalcontacts.

The paddle stimulation leads 141 consist of a combination of implantablematerials. Stimulation electrodes 130 at the distal end and electricalcontacts at the proximal end are made of a 90-10 platinum-iridium alloy.This alloy is utilized for its bio-compatibility and electricalconductivity. The polymeric body of the lead 141 is polyurethane, whichis chosen for its bio-compatibility, flexibility, and high lubricity todecrease friction while being passed through tissue. The polyurethanetubing has a multi-lumen cross section, with one center lumen 142 andeight outer lumens 144. The center lumen 142 acts as a canal to containthe steering stylet during implantation, while the outer lumens 144provide electrical and mechanical separation between the wires 146 thatcarry stimulation from the proximal contacts to distal electrodes 130.These wires 146 are a bundle of MP35N strands with a 28% silver core.The wires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. At the distaltip of the paddle leads 141, there is a 2-dimensional array of flatrectangular electrodes 130, molded into a flat silicone body 149. Onlyone side of the rectangular electrodes 130 is exposed, providing thedesired uni-directional stimulation. The wires 146 are laser welded tothe contacts and electrodes 130, creating an electrical connectionbetween respective contacts on the proximal and distal ends. Also moldedinto the distal silicone paddle is a polyester mesh 147 which addsstability to the molded body 149 while improving aesthetics by coveringwire 146 routing. The number of individual 8-contact leads 141 used foreach paddle 141 is governed by the number of electrodes 130. Electrodes130 per paddle 141 range from 8 to 32, which are split into between oneand four proximal lead 141 ends. Each proximal lead 141 has 8 contacts(+1 fixation contact). Configurations vary by electrode 130 number,electrode 130 spacing, electrode length, and overall lead length.

The lead extensions 150, as depicted in FIG. 10, are a fully implantableelectrical medical accessory to be used in conjunction with theimplantable SCS 100 and either percutaneous 140 or paddle 141 leads. Theprimary function of the lead extension 150 is to increase the overalllength of the lead 140, 141 by carrying electrical signals from the IPG102 to the proximal end of the stimulation lead 140, 141. This extendsthe overall range of the lead 140, 141 in cases where the length of theprovided leads 140, 141 is insufficient for case. The lead extensions150 must provide a robust, flexible, and bio-compatible electricconnection between the IPG 102 and proximal end of the stimulation lead140, 141. The extensions 150 may be secured mechanically to the patientusing either an anchor or a suture passed through tissue and tied aroundthe body of the extension 150. Extensions 150 are secured at theproximal end with a set-screw 119 on the IPG 102 which applies radialpressure to a fixation contact on the distal end of the proximalcontacts of the extension 150. The stimulation lead 140, 141 is securedto the extension 150 in a similar fashion, using a set screw 152 insidethe molded tip of extension 150 to apply a radial pressure to thefixation contact at the proximal end of the stimulation lead 140, 141.

The lead extension 150 consists of a combination of implantablematerials. At the distal tip of the extension 150 is a 1×8 array ofimplantable electrical contacts 154, each consisting of MP35 housing 128and 90-10 platinum-iridium spring. A silicone seal 127 separates each ofthe housings 128. At the proximal end of the contacts is a titanium(6Al4V) cap which acts as a stop for the lead, and at the distal tip, atitanium (6Al4V) block and set screw 152 for lead fixation. Theelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy. This alloy is utilized for its bio-compatibilityand electrical conductivity. The polymeric body 156 of the lead 150 ispolyurethane, which is chosen for its bio-compatibility, flexibility,and high lubricity to decrease friction while being passed throughtissue. The polyurethane tubing 158 has a multi-lumen cross section,with one center lumen 142 and eight outer lumens 144. The center lumen142 acts as a canal to contain the steering stylet during implantation,while the outer lumens 144 provide electrical and mechanical separationbetween the wires 146 that carry stimulation from the proximal contactsto distal electrodes. These wires 146 are a bundle of MP35N strands witha 28% silver core. The wires 146 are individually coated with ethylenetetrafluoroethylene (ETFE), to provide an additional non-conductivebarrier. Each lead extension 150 has 8 proximal cylindrical contacts (+1fixation contact).

The lead splitter 160, as depicted in FIG. 11, is a fully implantableelectrical medical accessory which is used in conjunction with the SCS100 and typically a pair of 4-contact percutaneous leads 140. Theprimary function of the lead splitter 160 is to split a single lead 140of eight contacts into a pair of 4 contact leads 140. The splitter 160carries electrical signals from the IPG 102 to the proximal end of two4-contact percutaneous stimulation leads 140. This allows the surgeonaccess to more stimulation areas by increasing the number of stimulationleads 140 available. The lead splitter 160 must provide a robust,flexible, and bio-compatible electrical connection between the IPG 102and proximal ends of the stimulation leads 140. The splitters 160 may besecured mechanically to the patient using either an anchor or a suturepassed through tissue and tied around the body of the splitter 160.Splitters 160 are secured at the proximal end with a set-screw 119 onthe IPG 102 which applies radial pressure to a fixation contact on thedistal end of the proximal contacts of the splitter 160. The stimulationleads 140 are secured to the splitter 160 in a similar fashion, using apair of set screws inside the molded tip of splitter 160 to apply aradial pressure to the fixation contact at the proximal end of eachstimulation lead 140.

The lead splitter 160 consists of a combination of implantablematerials. At the distal tip of the splitter 160 is a 2×4 array ofimplantable electrical contacts 162, with each contact 162 consisting ofMP35 housing 128 and 90-10 platinum-iridium spring. A silicone seal 127separates each of the housings 128. At the proximal end of each row ofcontacts 162 is a titanium (6Al4V) cap which acts as a stop for thelead, and at the distal tip, a titanium (6Al4V) block and set screw forlead fixation. The electrical contacts at the proximal end of thesplitter 160 are made of a 90-10 platinum-iridium alloy. This alloy isutilized for its bio-compatibility and electrical conductivity. Thepolymeric body 164 of the lead 160 is polyurethane, which is chosen forits bio-compatibility, flexibility, and high lubricity to decreasefriction while being passed through tissue. The polyurethane tubing 166has a multi-lumen cross section, with one center lumen 142 and eightouter lumens 144. The center lumen 142 acts as a canal to contain thesteering stylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. Each leadsplitter 160 has 8 proximal contacts (+1 fixation contact), and 2 rowsof 4 contacts 162 at the distal end.

Anchors

The lead anchor 170, as depicted in FIGS. 12 and 13, is a fullyimplantable electrical medical accessory which is used in conjunctionwith both percutaneous 140 and paddle 141 stimulation leads. The primaryfunction of the lead anchor 170 is to prevent migration of the distaltip of the lead 140, 141 by mechanically locking the lead 140, 141 tothe tissue. There are currently two types of anchors 170, a simplesleeve 171, depicted in FIG. 12, and a locking mechanism 172, depictedin FIG. 13, and each has a slightly different interface. For the simplesleeve type anchor 171, the lead 140, 141 is passed through the centerthru-hole 174 of the anchor 171, and then a suture is passed around theoutside of the anchor 171 and tightened to secure the lead 140, 141within the anchor 171. The anchor 171 can then be sutured to the fascia.The locking anchor 172 uses a set screw 176 for locking purposes, and abi-directional disposable torque wrench for locking and unlocking.Tactile and audible feedback is provided for both locking and unlocking.

Both anchors 171, 172 are molded from implant-grade silicone, but thelocking anchor 172 uses an internal titanium assembly for locking. The3-part mechanism is made of a housing 175, a locking set screw 176, anda blocking set screw 177 to prevent the locking set screw from back out.All three components are titanium (6Al4V). The bi-directional torquewrench has a plastic body and stainless steel hex shaft.

Wireless Dongle

The wireless dongle 200 is the hardware connection to asmartphone/mobile 202 or tablet 204 that allows communication betweenthe trial generator 107 or IPG 102 and the smartphone/mobile device 202or tablet 204, as illustrated in FIG. 14. During the trial or permanentimplant phases, the wireless dongle 200 is connected to the tablet 204through the tablet 204 specific connection pins and the clinicianprogrammer software on the tablet 204 is used to control the stimulationparameters. The commands from the clinician programmer software aretransferred to the wireless dongle 200 which is then transferred fromthe wireless dongle 200 using RF signals to the trial generator 107 orthe IPG 102. Once the parameters on the clinician programmers have beenset, the parameters are saved on the tablet 204 and transferred to thepatient programmer software on the smartphone/mobile device 202. Thewireless dongle 200 is composed of an antenna, a microcontroller (havingthe same specifications as the IPG 102 and trial generator 107), and apin connector to connect with the smartphone/mobile device 202 and thetablet 204.

Charger

The IPG 102 has a rechargeable Lithium ion battery 108 to power itsactivities. An external induction type charger 210 (FIG. 1) is necessaryto recharge the included battery 108 inside the IPG 102 wirelessly. Thecharger 210 consists of a rechargeable battery, a primary coil of wireand a printed circuit board (PCB) for the electronics—all packaged intoa housing. When switched on, this charger 210 produces magnetic fieldand induces voltage into the secondary coil 109 in the implant. Theinduced voltage is then rectified and then used to charge the battery108 inside the IPG 102. To maximize the coupling between the coils, bothinternal and external coils are combined with capacitors to make themresonate at a particular common frequency. The coil acting as aninductor L forms an LC resonance tank. The charger uses a Class-Eamplifier topology to produce the alternating current in the primarycoil around the resonant frequency. Below are the charger 210 features;

-   -   Charges IPG 102 wirelessly    -   Charges up to a maximum depth of 30 mm    -   Integrated alignment sensor helps align the charger with IPG 102        for higher power transfer efficiency    -   Alignment sensor gives an audible and visual feedback to the        user    -   Compact and Portable

A protected type of cylindrical Li ion battery is used as the charger210 battery. A Class-E of the topologies of the power amplifiers hasbeen the most preferred type of amplifier for induction chargers,especially for implantable electronic medical devices. It's relativelyhigh theoretical efficiency made it the most favorable choice fordevices where high efficiency power transfer is necessary. A 0.1 ohmhigh wattage resistor is used in series to sense the current throughthis circuit.

The primary coil L1 is made by 60 turns of Litz wire type 100/44-100strands of 44 AWG each. The Litz wire solves the problem of skin effectand keeps its impedance low at high frequencies. Inductance of this coilwas initially set at 181 uH, but backing it with a Ferrite plateincreases the inductance to 229.7 uH. The attached ferrite plate focusesthe produced magnetic field towards the direction of the implant. Such asetup helps the secondary coil receive more magnetic fields and aids itto induce higher power.

When the switch is ON, the resonance is at frequency

$f = \frac{1}{2\; \pi \sqrt{L\; 1C\; 2}}$

When the switch is OFF, it shifts to

$f = \frac{1}{2\; \pi \sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$

In a continuous operation the resonance frequency will be in the range

$\frac{1}{2\; \pi \sqrt{L\; 1C\; 2}} < f < \frac{1}{2\; \pi \sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$

To make the ON and OFF resonance frequencies closer, a relatively largervalue of C1 can be chosen by a simple criteria as follows

C1=nC2; a value of n=4 was used in the example above; in most cases3<n<10.

The voltages in these Class-E amplifiers typically go up to the order of300 VAC. Capacitors selected must be able to withstand these highvoltages, sustain high currents and still maintain low Effective SeriesResistance (ESR). Higher ESRs result in unnecessary power losses in theform of heat. The circuit is connected to the battery through aninductor which acts as a choke. The choke helps to smoothen the supplyto the circuit. The N Channel MOSFET acts as a switch in this Class-Epower amplifier. A FET with low ON resistance and with high draincurrent I_(d) is desirable.

In summary, the circuit is able to recharge the IPG 102 battery 108 from0 to 100% in 2 Hr 45 Min with distance between the coils being 29 mm.The primary coil and the Class-E amplifier draws DC current of 0.866 Ato achieve this task. To improve the efficiency of the circuit, afeedback closed loop control is implemented to reduce the losses. Thelosses are minimum when the MOSFET is switched ON and when the voltageon its drain side is close to zero.

The controller takes the outputs from operational amplifiers, checks ifthey meet the criteria, then it triggers the driver to switch ON theMOSFET for next cycle. The controller needs to use a delay timer, an ORgate and a 555 timer in monostable configuration to condition the signalfor driver. When the device is switched ON, the circuit does not startto function right away as there will be no active feedback loop. Thefeedback becomes active only if the circuit starts to function. To solvethis riddle, an initial external trigger is applied to jump start thesystem.

Alignment Sensor

The efficiency of the power transfer between the external charger 210and the internal IPG 102 will be maximum only when they are properlyaligned. An alignment sensor is absolutely necessary to ensure a properalignment. This is a part of the external circuit design. The firstdesign is based on the principle called reflected impedance. When theexternal is brought closer to the internal, the impedance of the bothcircuits change. The sensing is based on measuring the reflectedimpedance and test whether it crosses the threshold. A beeper is used togive an audible feedback to the patient; an LED is used for visualfeedback.

When the impedance of the circuit changes, the current passing throughit also changes. A high power 0.1 ohm resistor is used in the series ofthe circuit to monitor the change in current. The voltage drop acrossthe resistor is amplified 40 times and then compared to a fixedthreshold value using a operational amplifier voltage comparator. Theoutput was fed to a timer chip which in turn activates the beeper andLED to give feedback to the user.

This circuit was successfully implemented in the lab on the table topversion. The circuit was able to sense the alignment up to a distance of30 mm. The current fluctuation in the circuit depends on more factorsthan reflected impedance alone and the circuit is sensitive to otherparameters of the circuit as well. To reduce the sensitivity related toother parameters, one option is to eliminate interference of all theother factors and improve the functionality of the reflected impedancesensor—which is very challenging to implement within the limited spaceavailable for circuitry. Another option is to use a dedicated sensorchip to measure the reflected impedance.

A second design uses sensors designed for proximity detector or metaldetectors for alignment sensing. Chips designed to detect metal bodiesby the effect of Eddy currents on the HF losses of a coil can be usedfor this application. The TDE0160 is an example of such a chip.

The external charger is designed to work at 75 to 80 kHz, whereas theproximity sensor was designed for 1 MHz. The sensor circuit is designedto be compatible with the rest of the external and is fine tuned todetect the internal IPG 102 from a distance of 30 mm.

Programmer

The Clinician Programmer is an application that is installed on a tablet204. It is used by the clinician to set the stimulation parameters onthe trial generator 107 or IPG 102 during trial and permanentimplantation in the operating room. The clinician programmer is capableof saving multiple settings for multiple patients and can be used toadjust the stimulation parameters outside of the operations room. It iscapable of changing the stimulation parameters though the RF wirelessdongle 200 when the trial generator 107 or IPG 102 in the patient iswithin the RF range. In addition, it is also capable of setting orchanging the stimulation parameters on the trial generator 107 and/orthe IPG 102 through the internet when both the tablet 204 and thePatient Programmers on a smartphone/mobile device 202 both have accessto the internet.

The Patient Programmer is an application that is installed on asmartphone/mobile device 202. It is used by the patient to set thestimulation parameters on the trial generator 107 or IPG 102 after trialand permanent implantation outside the operating room. The clinicianprogrammer is capable of saving multiple settings for multiple patientsand can be transferred to the Patient Programmer wirelessly when theClinician Programmer tablet 204 and the Patient Programmersmartphone/mobile device 202 are within wireless range such as Bluetoothfrom each other. In the scenario where the Clinician Programmer tablet204 and the Patient Programmer smartphone/mobile device 202 are out ofwireless range from each other, the data can be transferred through theinternet where both devices 202, 204 have wireless access such as Wi-Fi.The Patient Programmer is capable of changing the stimulation parameterson the trial generator 107 or IPG 102 though the RF wireless dongle 200when the trial generator 107 or IPG in the patient is within the RFrange. However, the Patient Programmer has limitations to changing thestimulation parameters.

Tuohy Needle

The tuohy needle 240, as depicted in FIG. 15, is used in conjunctionwith a saline-loaded syringe for loss-of-resistance needle placement,and percutaneous stimulation leads 140, for lead 140 placement into thespinal canal. The tuohy epidural needle 240 is inserted slowly into thespinal canal using a loss-of-resistance technique to gauge needle 240depth. Once inserted to the appropriate depth, the percutaneousstimulation lead 140 is passed through the needle 240 and into thespinal canal.

The epidural needle 240 is a non-coring 14G stainless steel spinalneedle 240 and will be available in lengths of 5″ (127 mm) and 6″(152.4). The distal tip 242 of the needle 240 has a slight curve todirect the stimulation lead 140 into the spinal canal. The proximal end246 is a standard Leur-Lock connection 248.

Stylet

The stylet 250, as depicted in FIG. 16, is used to drive the tip of apercutaneous stimulation lead 140 to the desired stimulation zone byadding rigidity and steerability. The stylet 250 wire 252 passes throughthe center lumen 142 of the percutaneous lead 140 and stops at theblocking plug at the distal tip of the lead 140. The tip of the stylet250 comes with both straight and curved tips. A small handle 254 is usedat the proximal end of the stylet 250 to rotate the stylet 250 withinthe center lumen 142 to assist with driving. This handle 254 can beremoved and reattached allowing anchors 170 to pass over the lead 140while the stylet 250 is still in place. The stylet 250 wire 252 is aPTFE coated stainless steel wire and the handle 254 is plastic.

Passing Elevator

The passing elevator 260, as depicted in FIG. 17, is used prior topaddle lead 141 placement to clear out tissue in the spinal canal andhelp the surgeon size the lead to the anatomy. The passing elevator 260provides a flexible paddle-shaped tip 262 to clear the spinal canal ofobstructions. The flexible tip is attached to a surgical handle 264.

The passing elevator 260 is a one-piece disposable plastic instrumentmade of a flexible high strength material with high lubricity. Theflexibility allows the instrument to easily conform to the angle of thespinal canal and the lubricity allows the instrument to easily passthrough tissue.

Tunneling Tool

The tunneling tool 270, as depicted in FIG. 18, is used to provide asubcutaneous canal to pass stimulation leads 140 from the entrance pointinto the spinal canal to the IPG implantation site. The tunneling tool270 is a long skewer-shaped tool with a ringlet handle 272 at theproximal end 274. The tool 270 is covered by a plastic sheath 276 with atapered tip 278 which allows the tool 270 to easily pass through tissue.Once the IPG 102 implantation zone is bridge to the lead 140 entrancepoint into the spinal canal, the inner core 275 is removed, leaving thesheath 276 behind. The leads 140 can then be passed through the sheath276 to the IPG 102 implantation site. The tunneling tool 270 is oftenbent to assist in steering through the tissue.

The tunneling tool 270 is made of a 304 stainless steel core with afluorinated ethylene propylene (FEP) sheath 276. The 304 stainless steelis used for its strength and ductility during bending, and the sheath276 is used for its strength and lubricity.

Torque Wrench

The torque wrench 280, as depicted in FIG. 19, is used in conjunctionwith the IPG 102, lead extension 150 and lead splitter 160 to tightenthe internal set screw 119, which provides a radial force against thefixation contact of the stimulation leads 140, 141, preventing the leads140, 141 from detaching. The torque wrench 280 is also used to lock andunlock the anchor 170. The torque wrench 280 is a small, disposable,medical instrument that is used in every SCS 100 case. The torque wrench280 provides audible and tactile feedback to the surgeon that the lead140, 141 is secured to the IPG 102, extension 150, or splitter 160, orthat the anchor 170 is in the locked or unlocked position.

The torque wrench 280 is a 0.9 mm stainless steel hex shaft 282assembled with a plastic body 284. The wrench's 280 torque rating isbi-directional, primarily to provide feedback that the anchor 170 iseither locked or unlocked. The torque rating allows firm fixation of theset screws 119, 152 against the stimulation leads 140, 141 withoutover-tightening.

Trial Patch

The trial patch is used in conjunction with the trialing pulse generator107 to provide a clean, ergonomic protective cover of the stimulationlead 140, 141 entrance point in the spinal canal. The patch is alsointended to cover and contain the trial generator 107. The patch is alarge, adhesive bandage that is applied to the patient post-operativelyduring the trialing stage. The patch completely covers the leads 140,141 and generator 107, and fixates to the patient with anti-microbialadhesive.

The patch is a watertight, 150 mm×250 mm anti-microbial adhesive patch.The watertight patch allows patients to shower during the trialingperiod, and the anti-microbial adhesive decreases the risk of infection.The patch will be made of polyethylene, silicone, urethane, acrylate,and rayon.

Magnetic Switch

The Magnetic magnetic switch is a magnet the size of a coin that, whenplaced near the IPG 102, can switch it on or off. The direction themagnet is facing the IPG 102 determines if the magnetic switch isswitching the IPG 102 on or off.

1. A spinal cord stimulation device (100) comprising: an implantablepulse generator (102); a plurality of implantable stimulation electrodes(130); a plurality of implantable leads (140, 141) electricallyconnecting the implantable pulse generator (102) and the optional trialimplantable pulse generator (107) to the plurality of stimulationelectrodes (130); an external charger (210) having a primary chargingcoil; a clinician programmer application provided on a computing device(204); an optional patient programmer application provided on a mobiledevice (202); and a communication device (200) allowing communicationbetween the computing device (204) and the optional mobile device (202)and the implantable pulse generator (102) and the optional trialimplantable pulse generator (107).
 2. The spinal cord stimulation device(100) of claim 1, wherein the communication device (200) comprises awireless dongle.
 3. The spinal cord stimulation device (100) of claim 1,the implantable pulse generator (102) comprising: a casing (120) havingan epoxy header (114); a lead contact assembly (116); a circuit board(105) comprising a plurality of output capacitors (112), an applicationspecific integrated circuit (106) and a microcontroller (104), themicrocontroller (104) in communication with the communication device(200) and with the application specific integrated circuit (106); arechargeable battery (108); a feedthrough (122) further comprising pins(124) that connect to the circuit board (105); a RF antenna (110); and asecondary charging coil (109).
 4. The spinal cord stimulation device(100) of claim 1, wherein the communication device (200) is configuredto operate in the 402 MHz to 405 MHz range utilizing up to 10 channelsfor telemetry.
 5. The spinal cord stimulation device (100) of claim 3,the rechargeable battery (108) comprised of an implantable grade lithiumion battery having zero-volt technology wherein the battery (108) isconfigured to be inductively charged via the external charger (210). 6.The spinal cord stimulation device (100) of claim 1, wherein theplurality of implantable stimulation electrodes (130) are grouped instimulation sets, each stimulation set programmable with differentstimulation parameters.
 7. The spinal cord stimulation device (100) ofclaim 3, the application specific integrated circuit (106) comprising adigital section and an analog section, wherein the digital sectioncomprises digital elements, timing generators, a plurality ofcomparators, arbitration control, pulse burst conditioner, and electrodelogic, and wherein the analog section comprises field effect transistorsand a plurality of digital-to-analog convertors.
 8. The spinal cordstimulation device (100) of claim 7, wherein the pulse burst conditioneroutputs a bursted bi-phasic pulse train and the timing generatorgenerates rising and falling edge triggers for each phase of thebi-phasic pulse.
 9. The spinal cord stimulation device (100) of claim 7,wherein the arbitration control analyzes the timing generator envelopesignals and permits only one signal to be active at a time.
 10. Thespinal cord stimulation device (100) of claim 8, wherein the analogsection is configured to analyze the bi-phasic pulse and convert it toan analog signal outputted to one of the plurality of electrodes (130).11. The spinal cord stimulation device (100) of claim 1, wherein theclinician programmer application provided on a computing device (204) isconfigured to perform all data processing functions and transmitoperational data only to the microcontroller (104).
 12. The spinal cordstimulation device (100) of claim 5, wherein a voltage doubler full waverectifier converts an induced AC voltage into usable DC voltage.
 13. Thespinal cord stimulation device (100) of claim 5, wherein the secondarycharging coil (109) and the primary charging coil are combined withcapacitors so that the primary and secondary charging coils resonant ata common frequency.
 14. The spinal cord stimulation device (100) ofclaim 5, further comprising an alignment sensor to denote when theprimary charging coil and the secondary charging coil (109) are properlyaligned, wherein the alignment sensor is configured to provide feedbackto indicate alignment.
 15. The spinal cord stimulation device (100) ofclaim 3, wherein the lead contact assembly (116) comprises: a pluralityof locking housings (118) each locking housing (118) defining anaperture, the aperture configured to accept a silicone tube (123) andthe silicone tube (123) configured to accept the lead (140), and whereinthe locking housing (118) further defines a second aperture configuredto accept a compression lead lock mechanism (119); and a plurality oflead contacts (126), each lead contact (126) separated from an adjacentlead contact (126) by a silicone seal (127), wherein the lead contacts(126) are configured in stacks of up to 8 lead contacts (126) and eachstack is connected at one end to one locking housing (118) and whereinthe lead contacts (126) connect to the feedthrough (122).
 16. The spinalcord stimulation device (100) of claim 1, wherein stimulation parametersare input to the computing device (204) and the optional mobile device(202), said stimulation parameters transmitted to the IPG (102) via thecommunication device (200).
 17. The spinal cord stimulation device (100)of claim 1, further comprising a trial implantable pulse generator(107).
 18. A method of providing stimulation to a spinal cordcomprising: implanting an implantable pulse generator (102)subcutaneously; implanting a plurality of implantable stimulationelectrodes (130) in the epidural space; electrically connecting theplurality of implantable stimulation electrodes (130) to the implantablepulse generator (102) via a plurality of implantable leads (140, 141);inputting stimulation parameters on a computing device (204) via aclinician programmer application; communicating the stimulationparameters to the implantable pulse generator (102) via a wirelesscommunication device (200); and delivering electrical pulses based onthe stimulation parameters to the electrodes (130) where the pulses arebi-phasic.